1. Field of the Invention
This application is generally directed to a positron emission tomography (PET) scanner system that reconstructs an image based on acquired PET information, acquired position information of a patient bed, and acquired position of a detector system.
2. Discussion of the Background
The use of positron emission tomography (PET) is growing in the field of medical imaging. In PET imaging, a radiopharmaceutical agent is introduced into the object to be imaged via injection, inhalation, or ingestion. After administration of the radiopharmaceutical, the physical and bio-molecular properties of the agent will cause it to concentrate at specific locations in the human body. The actual spatial distribution of the agent, the intensity of the region of accumulation of the agent, and the kinetics of the process from administration to eventually elimination are all factors that may have clinical significance. During this process, a positron emitter attached to the radiopharmaceutical agent will emit positrons according to the physical properties of the isotope, such as half-life, branching ratio, etc.
The radionuclide emits positrons, and when an emitted positron collides with an electron, an annihilation event occurs, wherein the positron and electron are destroyed. Most of the time, an annihilation event produces two gamma rays at 511 keV traveling at substantially 180 degrees apart.
By detecting the two gamma rays, and drawing a line between their locations, i.e., the line-of-response (LOR), one can retrieve the likely location of the original disintegration. While this process will only identify a line of possible interaction, by accumulating a large number of those lines, and through a tomographic reconstruction process, the original distribution can be estimated. In addition to the location of the two scintillation events, if accurate timing (within few hundred picoseconds) is available, a time-of-flight (TOF) calculation can add more information regarding the likely position of the event along the line. Limitations in the timing resolution of the scanner will determine the accuracy of the positioning along this line. Limitations in the determination of the location of the original scintillation events will determine the ultimate spatial resolution of the scanner, while the specific characteristics of the isotope (e.g., energy of the positron) will also contribute (via positron range and co-linearity of the two gamma rays) to the determination of the spatial resolution the specific agent.
The collection of a large number of events creates the necessary information for an image of an object to be estimated through tomographic reconstruction. Two detected events occurring at substantially the same time at corresponding detector elements form a line-of-response that can be histogrammed according to their geometric attributes to define projections, or sinograms to be reconstructed. Events can also be added to the image individually.
The fundamental element of the data collection and image reconstruction is therefore the LOR, which is the line traversing the system-patient aperture. Additional information can be obtained regarding the location of the event. First, it is known that, through sampling and reconstruction, the ability of the system to reconstruct or position a point is not space-invariant across the field of view, but is better in the center, slowly degrading toward the periphery. A point-spread-function (PSF) is typically used to characterize this behavior. Tools have been developed to incorporate the PSF into the reconstruction process. Second, the time-of-flight, or time differential between the arrival of the gamma ray on each detector involved in the detection of the pair, can be used to determine where along the LOR the event is more likely to have occurred.
The above described detection process must be repeated for a large number of annihilation events. While each imaging case must be analyzed to determine how many counts (i.e., paired events) are required to support the imaging task, current practice dictates that a typical 100-cm long, FDG (fluoro-deoxyglucose) study will need to accumulate several hundred million counts. The time required to accumulate this number of counts is determined by the injected dose of the agent and the sensitivity and counting capacity of the scanner.
PET imaging systems use detectors positioned across from one another to detect the gamma rays emitting from the object. Typically a ring of detectors is used in order to detect gamma rays coming from each angle. Thus, a PET scanner is typically substantially cylindrical to be able to capture as much radiation as possible, which should be, by definition, isotropic. The use of partial rings and rotation of the detector to capture missing angles is also possible, but these approaches have severe consequences for the overall sensitivity of the scanner. In a cylindrical geometry, in which all gamma rays included in a plane have a chance to interact with the detector, an increase in the axial dimension has a very beneficial effect on the sensitivity or ability to capture the radiation. Thus, the best design is that of a sphere, in which all gamma rays have the opportunity to be detected. Of course, for application to humans, the spherical design would have to be very large and thus very expensive. Accordingly, a cylindrical geometry, with the axial extent of the detector being a variable, is realistically the starting point of the design of a modern PET scanner.
PET scanners have progressively increased the quality (density) of sampling by putting smaller and smaller crystals as elements in the PET detector ring. PET scanners have become more and more three-dimensional by adding more and more rings to the PET detector system. PET detector crystals are already one the most expensive portions of the PET scanner. Small animal scanners are pushing the limit of how small the scintillation crystals can be. While typical, human-sized clinical systems are built around 4-6 mm square or rectangular crystals, some small-animal systems are using effectively a sub-millimeter crystal size. However, the increased cost to prepare and characterize each crystals, and the complexity (smaller crystals also means more channels and more accurate reconstruction requirements), has put a heavy burden on the use of smaller crystals in clinical systems.
Another way to improve the spatial resolution of the acquired PET data is to perform sub-sampling. In this approach, a data set is acquired with the “native” geometry of the scanner, a second data set is then acquired with the geometry slightly changed. If the knowledge of the small change in geometry is more accurate than the native spatial resolution of the system, sub-sampling is then obtained with improvement in the overall system resolution. This approach with two samples can be repeated with more elements, but is limited to “single-ray” image reconstruction, such as filtered backprojection or Fourier rebinning, since modern statistical reconstruction methods (e.g., maximum-likelihood, expectation maximization (ML-EM)) do not benefit from such sampling as they already model the finite size of the sampling element.
Historically, two types of sub-sampling have been implemented for single-ray reconstruction systems. In one approach, a motion is imposed on the entire scanner in the transaxial plane (wobble), allowing for a smaller-than-crystal size resolution in the transaxial reconstruction. A second technique, very similar to what is also available in SPECT imaging, requires continuously moving the patient bed. In this approach, sub-crystal resolution can be obtained in the axial direction.
As shown in FIGS. 1A-1C, simple rotation and translation can drastically increase the density of sampling (line of response) from the same detection. FIGS. 1A-1C show sampling in the transaxial plane, including an original sampling pattern (FIG. 1A), a pattern obtained by a pure rotation about the main axis (FIG. 1B), and a pattern obtained by an up-down translation in the transaxial plane (FIG. 1C). FIGS. 2A and 2B show an original sampling pattern in the axial plane, and a pattern obtained by a simple lateral translation in the axial direction.